Imaging systems are increasingly important in medical technology. For example, imaging systems may be used to generate two- or three-dimensional image data of organs and structures of the human body. Such image data may be used, for example, to diagnose causes of illness, perform operations, and prepare therapeutic measures. The image data may be generated on the basis of measurement signals obtained with the aid of a radiation detector.
For example, in X-ray and computed tomography (CT) systems, the body or a body section of a patient to be examined is radiographed by X-ray radiation generated by a radiation source. The non-absorbed, transmitted portion of the radiation is detected by a detector.
By way of further example, in positron emission tomography (PET) systems and single photon emission computer tomography (SPECT) systems, image generation is achieved using radionuclides. The patient to be examined is injected with a radiopharmaceutical that generates gamma quanta either directly (e.g., in the case of SPECT) or indirectly (e.g., in the case of PET) through emission of positrons. The gamma radiation is detected by a corresponding radiation detector.
Detectors used for the energy-resolved detection or “counting” of radiation quanta may operate according to different measurement principles. For example, radiation may be detected either directly (e.g., by direct conversion of the radiation energy into electrical energy) or indirectly. In the case of indirect detection, a scintillator may be used. A scintillator is excited in response to the action of radiation to be detected and reemits the excitation energy by emitting lower-energy electromagnetic radiation. Only the radiation emitted by the scintillator is converted into electrical measurement signals. Detectors of planar construction (e.g., “flat detectors”) are used in the medical field and operate in accordance with these measurement principles (as described, for example, in: M. Spahn, “Flat detectors and their clinical applications,” Eur. Radiol., 2005, 15, 1934-1947).
The conversion of the radiation emerging from a scintillator into an electrical signal may be effected in various ways. For example, a photomultiplier provided with a photocathode in the form of an evacuated electron tube may be used. Alternatively, a silicon photomultiplier (SiPM) that involves a matrix arrangement of avalanche photodiodes (APD) embodied on a shared substrate may be used. The electrons are generated in the avalanche photodiodes as a result of incident photons, and are multiplied in an avalanche-like manner.
A disadvantage of silicon photomultipliers is that only part of the total area available for irradiation may be utilized as a sensitive or “active” area since there are insensitive regions (e.g., where resistors and signal lines or wiring structures are arranged) between the active or radiation-sensitive regions. Thus, a silicon photomultiplier has a relatively small ratio of active area to (irradiated) total area. This ratio is designated as a “filling factor”. Further disadvantages include the occurrence of noise during an operation and a relatively high dark rate or dark count (e.g., signal generation taking place even without irradiation).
In a detector having a scintillator and a silicon photomultiplier, the silicon photomultiplier may be opposite an end face or rear side of the scintillator, such that an opposite end face or front side of the scintillator faces the radiation to be detected. As a result, the silicon photomultiplier may detect only that portion of the radiation converted in the scintillator that emerges at the rear side thereof. However, the scintillation radiation emanating from the respective excitation or interaction location in the scintillator is emitted in other directions besides the direction of the rear side. Furthermore, the radiation is subject to loss processes (e.g., reflection, absorption, and scattering). The losses are relatively high for scintillators having a high aspect ratio (e.g., of height to width) as, for example, in a PET system. For an aspect ratio of greater than 7:1, the radiation emerging from a scintillator may constitute merely 40-60% of the total radiation generated. Although a higher intensity of the incident radiation may be provided to compensate for the loss, a patient would be exposed to an increased radiation dose as a result.
A further disadvantage of conventional detector design is that the interaction location of incident radiation in a scintillator is either undetectable or may be detected only with great difficulty based on the radiation emerging at the rear side of the scintillator. Moreover, information about the height or depth of an interaction in the scintillator may not be obtainable. Consequently, the resolution of imaging systems provided with such detectors is restricted.
Microchannel plates (MCP) having a multiplicity of channels have been used for image intensification and electron multiplication. During operation, an electrical voltage present along the channels is generated. As a result of the voltage, entering electrons may be accelerated within the channels and multiplied by impacts with the channel walls. The use of a microchannel plate in connection with an image intensifier is described in U.S. Patent Application Publication No. 2009/0256063 A1.
A PET detector module has been described by Wu Gao et al. [“Design of a Monolithic Multichannel Front-End Readout ASIC for PET Imaging Based on Scintillation Crystals Read Out by Photodetectors at Both Ends,” Advancements in Nuclear Instrumentation Measurement Methods and their Applications (ANIMMA), 2011 2nd International Conference on, IEEE, Jun. 6, 2011 (2011-06-06), pages 1-8, XP032153585, DOI: 10.1109/ANIMMA.2011.6172963, ISBN: 978-1-4577-0925-8]. The detector module includes scintillator crystals arranged alongside one another. A microchannel plate photomultiplier for detecting scintillation radiation is arranged at two opposite sides of the scintillator arrangement.
A further PET detector module is described by Heejong Kim et al. [“Continuous Scintillator Slab with Microchannel Plate PMT for PET,” 2009 IEEE Nuclear Science Symposium and Medical Imaging Conference (NSS/MIC 2009), Orlando, Fl., USA, IEEE, Piscataway, N.J., USA, Oct. 24, 2009 (2009-10-24), pages 2553-2556, XP031826651, DOI: 10.1109/NSSMIC.2009.5402028, ISBN: 978-1-4244-3961-4]. The detector module includes one or two layers of a plate-shaped scintillator. Microchannel plate photomultipliers for detecting scintillation radiation are arranged at the sides of the scintillators.
A SPECT camera including a scintillator, a photocathode, a microchannel plate, and a detection device having a photoanode is described in International Publication No. WO 2010/094272 A1.